Clinical Biomechanics 15 (2000) 352±358
www.elsevier.com/locate/clinbiomech
Biomechanical adaptations of transtibial amputee sprinting in
athletes using dedicated prostheses
John G. Buckley
Biomechanics Research Group, Department of Exercise and Sport Science, The Manchester Metropolitan University, Alsagar Campus, Hassell Road,
Stoke-on-Trent Cheshire ST7 2HL, Manchester, UK
Received 5 October 1999; accepted 01 December 1999
Abstract
Objective. To determine the biomechanical adaptations of the prosthetic and sound limbs in two of the worldÕs best transtibial
amputee athletes whilst sprinting.
Design. Case study design, repeated measures.
Background. Using dedicated sprint prostheses transtibial amputees have run the 100 m in a little over 11 s. Lower-limb
biomechanics when using such prostheses have not previously been investigated.
Methods. Moments, muscle powers and the mechanical work done at the joints of the prosthetic and sound limbs were calculated
as subjects performed repeated maximal sprint trials using a Sprint Flex or Cheetah prosthesis.
Results. An increased hip extension moment on the prosthetic limb, with an accompanying increase in the amount of concentric
work done, was the most notable adaptation in Subject 1 using either prosthesis. In Subject 2, an increased extension moment at the
residual knee, and an accompanying increase in the amount of total work done, was the most notable adaptation using either
prosthesis. This later adaptation was also evident in Subject 1 when using his Sprint Flex prosthesis.
Conclusions. Increased hip work on the prosthetic limb has previously been shown to be the major compensatory mechanism that
allow transtibial amputees to run. The increased work found at the residual knee, suggests that the two amputee sprinters used an
additional compensatory mechanism.
Relevance
These ®ndings provide an insight into the biomechanical adaptations that allow a transtibial amputee to attain the speeds
achieved when sprinting. Ó 2000 Elsevier Science Ltd. All rights reserved.
Keywords: Joint moments; Muscle powers; Amputee sprinting
1. Introduction
Below-knee amputees have run the 100 m in a little
over 11 s. At this level of performance, amputee
sprinting is a highly competitive sport, with many
athletes training as much as their able-bodied counterparts. Apart from developing the necessary motor
co-ordination, sprint training is undertaken by both
able-bodied and amputee athletes to develop strength
and power in the muscles of the lower-limb. Because
below-knee amputees require the use of an arti®cial
foot and ankle, both the bene®t of training and sprint
E-mail address: [email protected] (J.G. Buckley).
performance are a€ected by the design of their chosen
prosthesis.
Prostheses speci®cally designed for sprinting use
carbon ®bre materials to provide a ¯exible shank/foot
keel, which is able to deform elastically on loading and
recoil at toe-o€. Buckley [1] noted that below-knee
amputee athletes wearing Flex-Foot Modular III prostheses were able to achieve an Ôup-on-the-toesÕ gait
typical of able-bodied sprinting, and this resulted in
prosthetic limb kinematics which were similar to those
of the sound limb and able-bodied controls. Although
such an analysis provides a descriptive characterisation
of amputee sprinting, it gains only limited insight into
the dynamic function of these prostheses and provides
no information on what biomechanical adaptations are
required by the muscles and joints of the lower-limbs.
0268-0033/00/$ - see front matter Ó 2000 Elsevier Science Ltd. All rights reserved.
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J.G. Buckley / Clinical Biomechanics 15 (2000) 352±358
Such insight can be gained by determining the joint
moment and muscle power patterns of each limb. This
will not only indicate which generic muscle groups (e.g.,
¯exors or extensors) are active, but can indicate when
they are acting eccentrically to absorb power or concentrically to generate power.
This type of kinetic analysis has been undertaken in
below-knee amputee running of moderate pace. Miller
[2] reported an increased hip extension moment during
stance of the prosthetic limb, in subjects running at a
range of speeds (2.7±5.7 m sÿ1 ). Czerniecki and Gitter [3]
found, as would be expected, minimal power generation
and absorption for the prosthetic ankle in subjects
wearing a SACH foot whilst running at a velocity of 2.8
m sÿ1 . They concluded that increased stance phase hip
work on the prosthetic limb, and increased hip and knee
work on the intact limb during swing, were the major
compensatory mechanisms which allowed amputees to
run. Furthermore, Sanderson and Martin [4] were able
to show that amputees increased their running speed
from 2.7±3.5 m sÿ1 , by modulating the magnitude of the
moment developed at each joint, without altering the
temporal sequencing of these moments. It cannot however, be assumed that these biomechanical adaptations
found for amputee running of moderate pace will also
be present in amputee sprinting.
To the authorÕs knowledge, previous studies concerned with amputee running have not reported the use
of dedicated sports prostheses in their subjects. As
sprinting is the only mode of gait that is truly digitgrade,
a dedicated sports prosthesis is essential. Typically such
prostheses have no heel component and are permanently
set (in an unloaded state) in plantar¯exion so its length
is the same as the sound limb when standing on ÔtiptoesÕ. By determining the joint moments, muscle powers
and the work done by the musculature at the ankle, knee
and hip of the prosthetic and sound limbs, the purpose
of this study was to determine the biomechanical adaptations in the two best transtibial amputee sprinters in
the country. A subsequent aim was to examine the e€ect
of dedicated prosthetic design upon these adaptations
by comparing prosthetic limb kinetics determined when
subjects used a Sprint Flex and Cheetah prosthesis.
353
time was 13.3 s. The study met with local bioethics
committee approval and written informed consent was
obtained from both subjects.
Each subject used two dedicated sprint prostheses
during the study; one incorporating a Sprint Flex
Modular III (ÔSprint FlexÕ) prosthetic foot; the other a
Sprint Flex Modular IV (ÔCheetahÕ) prosthetic foot. Fig.
1 highlights the di€erent shape of each footÕs carbon®bre keel. Each foot was ®tted to its own patellar-tendon-bearing suction socket so its length was the same as
the sound limb when standing on tip-toe. Neither foot
was ®tted with a heel (keel) component. A spike plate
was ®tted to the bottom of the prosthetic foot (Fig. 1),
thus subjects wore a spiked running shoe on only their
sound foot. Both subjects had regularly used each of
their prostheses in training and in competition. The two
subjects were, at the time, the only two amputees in the
country to have such prostheses. The residual knee of
Subject 2 was unable to ¯ex beyond 90°, because it had
been pinned after the trauma that led to the transtibial
amputation.
2.2. Data collection
Data were collected over two testing sessions. During
the ®rst, subjects used their Sprint Flex prosthesis, and
in the second (which occurred over two months later)
their Cheetah prosthesis. After warming up, repeated
maximal sprint trials were performed along a 35 m indoor track. Starting positions, approximately 14 m from
a force platform, were adjusted to ensure either the
prosthetic foot or the intact foot landed squarely on the
platform. Data were only recorded if this was achieved
without either under- or over-striding. The platform
2. Methods
2.1. Subjects
Two transtibial amputee athletes volunteered to
participate. Both subjects had competed at all levels,
including World and Paralympian Games, and undertook regular athletic training on a weekly basis. Subject
1 (age, 25 yr; body mass, 72.6 kg; stature, 1.88 m) had a
personal best time for the 100 m sprint of 12.7 s. Subject
2Õs (age, 24 yr; body mass, 75 kg; stature, 1.78 m) best
Fig. 1. The di€erent shapes of the Sprint Flex and Cheetah prosthesis
carbon-®bre keel are shown. Also shown are the position of markers
used in modelling the prosthesis. Markers were placed on the prosthetic socket, at a position corresponding to the underlying knee
centre, on the carbon-®bre foot keel, either at the same height as the
lateral malleolus of the intact limb when standing on tip-toe (Sprint
Flex), or at the point were the radius was most acute (Cheetah), and on
the top surface of the keel, 2 cm proximal of the most distal point.
354
J.G. Buckley / Clinical Biomechanics 15 (2000) 352±358
surface, like the surrounding ¯oor, was covered with
0.014 m thick tartan covering to Altro Mondo Sport¯ex
speci®cation. The order in which each foot landed on
the platform was randomised, and trials were repeated,
giving sucient rest between each trial, until between 6
and 8 successful trials were completed for each limb.
Although the number of trials might seem high, each
subject was used to running multiple trials in training
and was able to complete the testing with relative ease.
An infra-red camera, mounted on a tripod, and positioned approximately 5.5 m from and perpendicular to
the plane of movement, tracked movements in the sagittal plane. Retro-re¯ective hemispherical markers (10
mm diameter) were attached to the following body
landmarks; anterior superior iliac spine (ASIS), greater
trochanter, lateral femoral condyle, lateral malleolus
and head of ®fth metatarsal. Markers were placed on the
prosthetic socket, at a position corresponding to the
underlying knee centre, on the carbon-®bre foot keel,
either at the same height as the lateral malleolus of the
intact limb when standing on tip-toe (Sprint Flex), or at
the point where the radius was most acute (Cheetah),
and on the top surface of the keel, 2 cm proximal of the
most distal point (Fig. 1). Calibration of the camera
system had previously been performed by recording the
position of a planar marker matrix (400 ´ 400 mm), of
length 2.0 m and height 1.2 m, which was placed centrally over the force platform. Ground reaction force
data and kinematic data were collected simultaneously
at 100 Hz using an Elite Motion Analysis system (Bioengineering Technology Systems, Italy), as subjects
sprinted over the platform. When using a 2D, single
camera protocol, body movements which occur outside
the movement plane (i.e., the plane perpendicular to the
optical axis of the camera) cannot be determined accurately because of problems associated with perspective
and/or parallax errors. However, as sprinting is achieved
by body movements predominantly in the sagittal plane,
such a protocol was considered sucient to meet the
aims of the present study. Kinematic data were
smoothed using the systemÕs automatic (lambda technique) ®ltering software [5]. Vertical and fore-and-aft
force, centre of pressure location, and the smoothed coordinate data of each marker were exported to ASCII
®le for further analysis.
To give an indication of the speed that each subject
sprinted over the force platform, the mean horizontal
velocity of the hip joint through the calibrated area (1 m
before and 1 m after the force platform) was determined
for each trial.
Subject 1 achieved a mean speed for the repeated
trials of 6.81 (S.D. 0.06) and 6.95 (S.D. 0.10) m sÿ1 using
the Cheetah and Sprint Flex prostheses, respectively,
whilst Subject 2 achieved speeds of 6.84 (S.D. 0.21) and
7.05 (S.D. 0.05) m sÿ1 . These speeds indicate that subjects were close to attaining their full sprinting speed.
2.3. Data analysis
Sagittal plane net joint moment and muscle power
outputs at the ankle, knee and hip joints were calculated
using a standard inverse dynamics approach [6]. All
body segment parameters were derived from the regression relationships reported by Drillis and Contini
[7]. In an attempt to evaluate how the mechanical behaviour of the prosthesis replicated that of the intact
limb (foot/ankle) it was modelled in a similar fashion,
i.e., it was modelled as comprising of a separate foot and
shank segment with simple hinge joint between the two.
It is recognised [8,9] that this approach is limited because, in addition to the underlying assumptions of the
linked segment model (that the body is a series of rigid
segments linked by frictionless joints), it assumes that
de¯ection of the prosthesis occurs about a ®xed centre of
rotation. However, as Czerniecki et al. [9] suggest,
``consistent marker placement on the prosthesis will result in reasonable approximations of the true power
output''. Such an approach also allows comparisons
between types of prosthetic feet. Miller [2] has previously demonstrated that compared to determining inertia characteristics experimentally, the use of standard
regression equations to determine the inertia characteristics of the prosthetic foot and shank, e€ected the resultant moment by less than 3%. This ®nding, combined
with the diculty of accurately assessing the inertia
characteristics of the below-knee stump and prosthesis,
as Miller [2] highlight, justi®es the use of standard regression equations to estimate the mass, centre of mass
and moment of inertia of both the prosthetic and intact
limb. This approach has previously been used to determine joint kinetics in amputee running [2±4,8,9].
Muscle power output was calculated as
Pj ˆ Mm xj
…W†;
where Pj is the joint muscle power, Mm the joint muscle
moment (N m), and xj is the joint angular velocity (rad
sÿ1 ).
Positive power resulted if the vectors Mm and xj had
the same polarity, i.e., either both had positive direction
or both had negative direction. This approach determines power for the ÔgenericÕ muscle groups, i.e., joint
¯exors or extensors. The work done (concentric or eccentric) by these generic muscles was calculated by integrating (numerically) the power output curve.
Z
Wmj ˆ Pj dt;
where Wmj is the generic muscle work, and Pj is the joint
power (instantaneous).
Where by the area under the power curve that was
negative re¯ected eccentric work whilst the area that was
positive re¯ected concentric work.
J.G. Buckley / Clinical Biomechanics 15 (2000) 352±358
The data (time series) were normalised to percentage
of each subjectÕs stance time. Results are presented as
the mean and S.D. of the repeated trials. Findings for
the ``Sprint Flex'' and ``Cheetah'' prosthetic limbs were
compared to the sound limb, and to each other.
3. Results
Results indicate that the joint moments and muscle
power outputs on the prosthetic limb were di€erent to
those determined for the sound side. There was also
evidence that prosthetic foot design had an in¯uence on
these ®ndings.
Apart from di€erences for the knee and hip when the
Cheetah prosthesis was used, ®ndings for Subject 2 were
similar to Subject 1. Thus, to avoid repetition, only
®ndings for Subject 1 are shown graphically. Figs. 2 and
3 show, respectively, the joint moment pattern and
muscle power output for the sound limb and each
prosthesis. As ®ndings for the sound limb showed little
355
di€erence when using either prosthesis, the data presented for the sound limb are from the ``Cheetah'' trials.
Peak joint moments and powers for the sound and each
prosthetic limb, of both subjects, are presented in Tables
1 and 2, respectively. And the joint work done (either
concentrically or eccentrically) by each limb is shown in
Fig. 4.
Each prosthesis produced an ÔankleÕ extension moment throughout the stance period. The power determined for the prosthetic ÔankleÕ showed that each
arti®cial foot, like the ankle of the sound limb, absorbed
power during the ®rst half of stance and generated
power during the second half. Peak power values,
however, were considerably less than those determined
for the sound limb (Table 2). The resulting values in
energy absorbed and returned by each prosthesis had
more-or-less the same magnitude, with the values for the
Sprint Flex prosthesis approximately 20 J higher than
those for the Cheetah prosthesis. On the sound limb
greater energy was generated at the ankle than absorbed.
Fig. 2. Joint moments of the ankle, knee and hip of the prosthetic (solid line) and sound limbs (dotted line) in Subject 1 wearing either the Sprint Flex
or Cheetah prosthesis. The mean is shown along with ‹1 S.D. for the prosthetic limb, but to aid clarity, only the mean is shown for the sound limb.
Moments that tended to extend a joint are shown as positive, and those that tended to ¯ex a joint are shown as negative.
356
J.G. Buckley / Clinical Biomechanics 15 (2000) 352±358
Fig. 3. Joint moments of the ankle, knee and hip of the prosthetic (solid line) and sound limbs (dotted line) in Subject 1 wearing either the Sprint Flex
or Cheetah prosthesis. The mean is shown along with ‹1 S.D. for the prosthetic limb, but to aid clarity, only the mean is shown for the sound limb.
Positive power indicates concentric action, and negative power indicates eccentric action.
When Subject 1 used his Sprint Flex prosthesis and
Subject 2 used either prosthesis, the residual knee produced an extension moment throughout stance. The
resulting power indicated that moderate eccentric work
was performed during the ®rst part of stance. Then
following midstance, considerable concentric work was
done accompanying a burst of power generation (Fig. 3).
When Subject 1 used his Cheetah prosthesis a ¯exion
moment was developed during the ®rst half of stance,
with only a moderate extension moment being developed during the second half (Table 1). The resulting
power was negligible throughout stance.
When Subject 2 used his Cheetah prosthesis, a hip
¯exion moment was evident on the prosthetic side
throughout stance and the corresponding power curve
indicated that only eccentric work (negative power) was
Table 1
Peak moments (N m) at the joints of the prosthetic and sound limbsa
Muscle involved
Ankle
Knee
Hip
a
Plantar¯exor
Dorsi¯exor
Flexor
Extensor (mid-stance)
Extensor
Flexor (late stance)
Subject 1
Subject 2
Sound
Sprint Flex
Cheetah
Sound
Sprint Flex
Cheetah
352
±
±
170
200
169
256
±
±
234
190
217
203
±
100
45
281
162
267
±
120
121
290
344
230
±
±
259
108
264
185 (29)
±
±
461 (94)
±
390 (143)
(28)
(59)
(84)
(43)
(21)
(26)
(32)
(52)
Values are the mean of the repeated trials and S.D. are shown in brackets.
(63)
(52)
(22)
(52)
(23)
(12)
(25)
(26)
(64)
(41)
(45)
(6)
(24)
(32)
J.G. Buckley / Clinical Biomechanics 15 (2000) 352±358
357
Table 2
Peak muscle powers (W) for the joints of the prosthetic and sound limbsa
Muscle action
Ankle
Knee
Hip
a
Ecc plantar¯exor
Con plantar¯exor
Ecc ¯exor
Ecc extensor
Con extensor
Con extensor
Ecc ¯exor
Subject 1
Subject 2
Sound
Sprint Flex
Cheetah
Sound
Sprint Flex
Cheetah
2336 (839)
2741 (467)
±
501 (352)
879 (382)
1349 (832)
1535 (405)
738 (171)
1012 (135)
±
542 (66)
1431 (55)
991 (344)
1714 (579)
681 (254)
307 (255)
Negligible
±
187 (105)
1891 (420)
1058 (149)
1822 (281)
1853 (238)
Negligible
±
861 (190)
2054 (428)
3090 (501)
1093 (312)
870 (271)
±
647 (245)
1416 (389)
571 (18)
1951 (745)
700 (171)
637 (259)
±
1469 (601)
2686 (742)
±
4437 (1739)
Values are the mean of the repeated trials and S.D. are shown in brackets.
Fig. 4. Mechanical work done (J) at the ankle (left column), knee
(middle column) and hip (right column) joints of the prosthetic and
sound limbs. The height of each column indicates the total work done,
the shaded portion represents the eccentric work and the un-shaded
portion the concentric work.
performed. In all other cases, the hip of the prosthetic
limb developed an extension moment which was maintained until late stance. The resulting power, showed
that this prolonged extension moment was created from
concentric action (positive power). On the sound side, a
short duration extensor moment was evident in early
stance, then a ¯exion moment predominated and was
associated with negative power.
4. Discussion
Previously, the only amputee athletic gait that had
been analysed biomechanically in detail was transtibial
amputee running of moderate pace. The present study
determined (stance phase) biomechanical adaptations in
two of the worldÕs best transtibial athletes whilst
sprinting using dedicated prostheses. In one of the
subjects (Subject 1) an increased and prolonged hip extension moment on the prosthetic limb, with an accompanying increase in the amount of concentric work
done was the most notable adaptation using either
prosthesis. This is in agreement with Czerniecki and
GitterÕs [3] suggestion, that ``the major compensatory
mechanism which allow transtibial amputees to run is
increased hip work on the prosthetic limb during
stance.'' However, in Subject 2 an increased extension
moment at the residual knee, and an accompanying increase in the amount of eccentric and concentric work
done by this joint was the most notable adaptation using
either prosthesis. It would be convenient to explain the
disparity in Subject 2Õs ®ndings and those previously
reported for running, by suggesting that sprinting requires completely di€erent biomechanical adaptations
than in running. However, the ®ndings for Subject 1 (of
increased hip work on the prosthetic limb) indicate this
is not the case. In addition, the increased moment and
mechanical work done at the residual knee (in Subject 2)
may have resulted because the knee of the intact limb
under-performed because it had been pinned following
the trauma which led to the transtibial amputation.
However, Subject 1 also had an increased extension
moment and an increase in the amount of work done at
the residual knee when using the Sprint Flex prosthesis.
It may well be that these kinetic di€erences, determined
for the residual knee, are additional, rather than completely di€erent, adaptations used by the two amputee
subjects to attain the speeds achieved when sprinting.
In previous studies concerned with amputee running,
no mention is made of the use of dedicated sports
prostheses in their subjects. In the present study subjects
used two types of dedicated sprint prosthesis, each incorporating a single continuous carbon ®bre (shank/
foot) keel. The shape of the Cheetah prosthesis is supposedly based on the shape of a hind leg of a cheetah
(i.e., like an up-side-down question mark), whilst the
Sprint Flex prosthesis is J-shaped. As sprinting is the
only mode of gait that is truly digitgrade, a dedicated
prosthesis is essential. In a previous study [1] in our
laboratory, it was noted that amputee athletes wearing
FlexFoot Modular III prostheses, which were set in
plantar¯exion so their length was the same as the sound
limb when standing on tip-toes, were able to achieve an
Ôup-on-the-toesÕ gait typical of able-bodied sprinting. In
the present study, subjects also achieved an Ôup-on-the-
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J.G. Buckley / Clinical Biomechanics 15 (2000) 352±358
toesÕ gait. At initial ground contact (which was made on
the ÔtoeÕ region of the prosthesis) the carbon ®bre keel
elastically deformed and consequently absorbed power.
Then following midstance power was generated as the
keel returned to its original (unloaded) shape. The
subsequent work done during each phase was approximately the same. Czerniecki et al. [9] studied the in¯uence of energy storing feet on the biomechanics of
transtibial amputee running. Using the same approach
to model the prosthesis, to that used in the present
study, they determined the power outputs of three different prosthetic feet. Their ®ndings indicate that when
subjects used a SACH foot, 31% of the energy absorbed
by the prosthesis was returned. In contrast 51% and 82%
was returned when using a Seattle and Flex foot, respectively. The ®ndings of the present study, indicate
that 100% of the energy absorbed by the prosthesis was
returned. The reduced number of prosthetic features of
these dedicated prostheses may explain this ecient return of energy. The prostheses used in Czerniecki et al.Õs
study [9], which were presumably aligned and ®tted for
everyday use, all incorporated a heel component. Such
prosthetic heel components are designed to deform at
ground contact and thereby reduced the loads transferred to the residual stump. The resulting energy absorbed at ground contact is typically used to initiate
dorsi¯exion of the foot and hence encourage the transition of the amputeeÕs weight onto the ÔforefootÕ section of
the prosthesis. Thus this energy is dissipated before toeo€. In the present study subjects landed on the ÔtoeÕ section of their (pre-plantar¯exed) prosthesis, and this was
also the point from which toe-o€ occurred. Thus the
energy absorbed (as the foot deformed) during initial
contact and the ®rst half of stance, could be directly returned during the second half of stance, as the keel
tended to return to its original shape prior to toe-o€. In
theory this spring-like action would only occur if the
frequency of the deformation and recoil of the prosthesis
was optimum [10]. The results presented here suggest this
was the case, although the Ôspring-likeÕ behaviour of such
prostheses certainly warrants further investigation.
Although there were di€erences in the joint kinetics of
each subject when using the two prostheses, there was
no obvious indication that the subjects favoured a particular type. Each subjectÕs personal rating of the two
prostheses corroborated this contention. Subject 1, indicated that he preferred the Sprint Flex prosthesis.
Whilst Subject 2 favoured the Cheetah prosthesis. The
total work done (either concentrically or eccentrically)
by the prosthetic limb when using each prosthesis would
tend to support these personal preferences. In Subject 1,
greater total work was done by the prosthetic limb when
using the Sprint Flex prosthesis compared to that done
when using the Cheetah prosthesis (S: 280.9 J; C: 180.1
J). Whilst Subject 2 did more total work on the prosthetic limb when using the Cheetah prosthesis (S: 225.2
J; C: 453.0 J). It may well be that the spring properties of
each subjectÕs ÔfavouredÕ prosthesis was (by chance)
better suited to his style of sprinting (e.g., ground contact duration). As more becomes known about the
biomechanical properties and in¯uence of such prostheses, it may be possible to tailor a prosthesis to optimally match an individualÕs requirements.
In the present study, joint kinetics were determined in
two of the worldÕs best transtibial amputee athletes,
during a single ground contact, whilst sprinting maximally. The ®ndings presented provide some insight, into
the biomechanical adaptations adopted during full
speed sprinting, as well as the mechanical behaviour of
the prostheses used. As sprinting requires an athlete to
accelerate from stationary up to their maximal speed in
the shortest amount of time, it would also be interesting
to investigate how such prostheses are utilised in the
early stages of a sprint race. It may well be that a
prosthesis, which facilitates full speed sprinting (because
its spring properties are optimum for this speed), behaves mechanically, quite di€erently during the initial
acceleration phase.
Acknowledgements
This work is part of an ongoing study, in collaboration with and partly funded by Chas A Blatchford and
Sons Ltd., Basingstoke, Hampshire, UK, investigating
the biomechanics of running and sprinting in lower-limb
amputee subjects.
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